JPH08131460A - Dental artificial tooth root and manufacturing method thereof - Google Patents

Abstract

(57) [Summary] (Modified) [Purpose] To provide a dental artificial tooth root having excellent bonding strength with a biocompatible material and a method for producing the same. [Structure] In a dental artificial dental root having a metal core material whose surface is coated with a biocompatible material, the metal core material is molded from an amorphous alloy having a supercooled liquid region, and titanium or titanium alloy is used for the metal core material. Compared with the conventional one,
It is possible to provide for the first time an artificial dental root having excellent bonding strength with a biocompatible material, and also when manufacturing an artificial dental root in which the metal core material is an amorphous alloy, a very simple processing step Since it can be manufactured with, it can contribute to cost reduction of the artificial tooth root itself.

Description

Detailed Description of the Invention

[0001]

BACKGROUND OF THE INVENTION The present invention relates to a dental artificial tooth root,
In particular, the present invention relates to an artificial tooth root having a metal core material whose surface is coated with a biocompatible material, and an improvement in a method for producing the artificial tooth root.

[0002]

2. Description of the Related Art Since an artificial dental root is to be implanted in the jawbone of a patient with an artificial dental crown, biocompatibility is strongly required in addition to mechanical strength. In recent years, it has been proposed to coat the surface of the metal core material with a biocompatible material such as hydroxyapatite. As one example, by plasma-spraying a biocompatible material on the surface of the metal core material, an artificial dental root having a coating layer of the biocompatible material formed on the surface of the metal core material has been developed. In that case, since the surface of the metal core material and the biocompatible material are bonded only by a chemical reaction, naturally,
There is a big problem that the joint strength between the two becomes weak and the coating layer of the biocompatible material is easily chipped or peeled off. In addition, once chipping or peeling occurred, the coating layer of the biocompatible material could be peeled off over a wide range.

Therefore, for example, Japanese Patent Laid-Open No. 6-205794.
An improved type has been developed as shown in the publication. This improved artificial dental root has particles of a biocompatible material attached to the surface of a metal core material made of titanium or a titanium alloy, and by utilizing the superplasticity of titanium, the particles of the biocompatible material of the metal core material Since the surface is embedded by hot pressing, the mechanically bonded state of the biocompatible material to the metal surface can be obtained.

[0004]

However, in such an improved artificial tooth root, titanium having 900 to 1000 is provided.
Since the superplastic region at ° C is used, 900
Since a relatively strong press pressure must be applied under a high temperature of ° C or higher, this time,
The particles of the biocompatible material may be crushed, and an unnecessary gap is created between the biocompatible material and the concave portion on the surface of the core material in which the biocompatible material is embedded, so that the holding force of the biocompatible material is increased. Problems such as weakening are newly introduced. Therefore, under the improved artificial dental root, even if mechanically bonding strength can be expected theoretically, in reality, further improvement is inevitable.

[0005]

In view of the above points, the present invention is premised on a dental artificial tooth root having a metal core material whose surface is coated with a biocompatible material. The present invention provides an artificial dental root made of an amorphous alloy having Therefore, in the dental dental implant according to the present invention, by using an amorphous alloy having a supercooled liquid region in the metal core material, as compared with the conventional one using titanium or titanium alloy in the metal core material, It is possible to provide an artificial dental root for which the mechanical bonding strength with the biocompatible material is further ensured. Also, especially,
When an amorphous alloy having a wide supercooled liquid region of 20 K or more or an amorphous alloy ratio of 30% or more is used, a high superplastic phenomenon that is not the ratio of titanium or titanium alloy is obtained in a predetermined relatively low temperature range. That is, the Ti-6Al-4VELI alloy is said to exhibit 1000% superplasticity at a temperature of 850 ° C.
Under the present invention, about 800 at 443K (170 ° C)
%, Some have a superplasticity of about 15000% at 473K (200 ° C).

The specific alloy composition is represented by the general formula:
Xa-Yb-Mc, where X is one or more metals selected from Zr, Ti, Hf, Mg and rare earth metals, and Y is Al, Z
One or more metals selected from r, Hf, Ti and rare earth metals, M is one or more metals selected from transition metals such as Fe, Co, Ni and Cu, a = 30 to 80, b = 5 20,
Amorphous alloys with c = 0-60 are possible. That is, since the base elements such as Zr, Ti, and Hf have almost no danger of being harmful to the human body, they are also suitable as a metal core material of a dental artificial dental root, and Fe of M element,
Co, Ni, and Cu coexist with the Zr, Ti, and Hf elements, etc., and improve the amorphous forming ability. Of these, a typical example that seems to be more preferable is Zr ₆₃ -A.
_{l 12 -Co 3 -Ni 7 -Cu 15} , Zr 60 -Al 15 -Co
_{5 -Ni 15 -Cu 5, Zr 65} -Al 7. 5 -Cu 27. 5,
_{Zr 55 -Al 20 -Co 20, Zr} 70 -Al 15 -Fe 15,
It is Zr ₆₀ -Al ₁₅ -Ni _25.

Further, in the case of producing a dental artificial dental root in which the metal core material is made of an amorphous alloy, a sintered body of a biocompatible material is formed into a cylindrical shape, and the inside of the cylindrical sintered body is molded. The amorphous alloy in the supercooled liquid region is pushed into, or while the particles of the biocompatible material are attached to the surface of the amorphous alloy core material in the supercooled liquid region, the particles of the biocompatible material are Only by embedding it in the surface, it becomes possible to easily manufacture a dental artificial dental root in which the metal core material is an amorphous alloy.

[0008]

DESCRIPTION OF THE PREFERRED EMBODIMENTS The present invention will be described below in detail based on embodiments.
First, as shown in FIG. 1, each mother alloy 1 having the above-mentioned alloy composition as a representative example is provided with φ1.0 at the tip.
Put it in the quartz nozzle 2 with a small hole 2a of ~ 2.0 mm,
After melting by induction heating in a vacuum, the quartz nozzle 2 is lowered, and then the pressure of argon gas is applied to the copper mold 3.
By injecting the molten metal into the inside from the small holes 2a, 1
Φ to be a sample by quenching at a cooling rate of 0 ⁴ to 10 ⁷ K / s
A 3.0 × 50 mm bulk material 4 was manufactured. The mother alloy 1 was made by melting sponge-like Zr metal in an arc melting furnace and degassing it, and then adding and melting other elements.

Then, when structural analysis was performed by an X-ray diffractometer to determine whether or not the bulk material 4 became an amorphous alloy, as shown in FIG. 2, the air-cooled material showed the presence of sharp Zr. The bulk material 4 made of No. 6 did not show a clear peak in the Kα ray equivalent region of Zr. Also, this amorphous alloy is preferably 30%
It has the above amorphous alloy ratio.

Next, differential scanning calorimetric analysis (DSC curve) was performed to examine the thermal properties of the amorphous alloys having the respective alloy compositions. As a result, as shown in FIG. 3, the glass transition temperature Tg and the crystallization temperature Tx differ greatly depending on the individual alloy composition, and the supercooled liquid region ΔTx
The maximum value of (Tx-Tg) is said to be 56K for the alloy composition containing Fe, 69K for the alloy composition containing Co, 77K for the alloy composition containing Ni, and 88K for the alloy composition containing Cu. The value was obtained. As a result, it was also found that the amorphous alloy having an alloy composition containing Cu showed the widest supercooled liquid region ΔTx, and Tg and Tx which gave this supercooled liquid region ΔTx were in a low temperature region of 630 to 710K. .

FIG. 4 shows the compositional dependence of the composition ratio of each of the three elements and the supercooled liquid region ΔTx for the amorphous alloy having the alloy composition containing Cu, which showed the best result in the DSC curve. Although it is shown, as is clear from the figure, 59 to 77 Zr, 5 to 13 Al, 25
A ～32Cu, particularly, the Zr ₆₅ -Al _{7. 5} -Cu
_27. The highest 80K supercooled liquid region of ₅ alloy composition △ T
x is shown. Although not shown, the alloy composition showed 100K and the alloy composition showed 90K.

FIG. 5 shows a stress-strain curve with respect to a temperature change in the vicinity of the supercooled liquid region ΔTx before Tg of the amorphous alloy having the alloy composition shown in FIG. It can be confirmed that the strain is caused by a slight stress. This indicates a superplastic phenomenon in the supercooled liquid region ΔTx. Also, 550 before Tg
It was found that even at K and 590K, the deformation started with a relatively weak stress, and thereafter the stress was relieved and the deformation proceeded continuously.

[0013] Further, Cu _{27 5} the Co, Ni, for others the alloy composition and was replaced by Cu, similar differential scanning calorimetry (DSC curve) and stress -. Investigated the relationship between strain curve and the temperature However, almost the same result was obtained.

Further, the amorphous alloy having the alloy composition of and has a Vickers hardness Hv450 and a tensile strength of 1400 MPa at room temperature and body temperature (37 ° C.),
The alloy composition (1) had a Vickers hardness of Hv400 and a tensile strength of 1200 MPa. These strengths are higher than those of pure titanium and titanium alloys such as Ti-6Al-4VELI.

Next, a method of manufacturing an artificial tooth root using the amorphous alloy having the above characteristics as a metal core material will be described. First, in the first manufacturing method, as shown in FIG. 6, the small holes 2a of the nozzle 2 are provided between rotating twin rolls 5 made of copper.
The molten metal is ejected from the molten metal to obtain an amorphous alloy wire 6 having a desired diameter, and the amorphous alloy wire 6 is used to form a metal core material having a predetermined size and shape. On the other hand, as shown in FIG. Outer diameter φ3.0-4.0 mm, height 10-1
A bottomed cylindrical sintered body 10 of hydroxyapatite, which is a biocompatible material, having a size of 3 mm was prepared.
On the other hand, as shown in FIG. 8, a temperature atmosphere having a glass transition temperature Tg or higher and a crystallization temperature Tx or lower (about 400 to 500 ° C.) is used.
In C), the amorphous alloy core material 12 in the supercooled liquid region is forcedly pushed. The particle size of the hydroxyapatite is preferably in the range of 10 to 300 μm.

Then, the amorphous alloy has a glass transition temperature of Tg or higher and a crystallization temperature of Tx or lower in a range of 200.
Since it exhibits 0% superplasticity, the amorphous alloy core material 12 surely enters into each of the micropores of the apatite sintered body 10, whereby the surface of the amorphous alloy core material 12 is apatite sintered. The body 10 will be joined with necessary and sufficient mechanical strength. Therefore, since the artificial tooth root manufactured by such a manufacturing method does not need to be subjected to hot pressing as in the conventional case, particles of hydroxyapatite are crushed during pressurization, and hydroxyapatite and it are embedded. There is no need to worry that an unnecessary gap is generated between the core material and the concave portion on the surface of the core material, and the holding force of the apatite coating layer is weakened.

If a heated molten metal of titanium or titanium alloy is poured into the apatite sintered body 10, the sintered body 10 may be damaged by this pressure, and as the molten metal cools and solidifies, There is a sufficient risk of the apatite sintered body 10 peeling or cracking due to the difference in thermal expansion coefficient from the hydroxyapatite, but there is no such concern under the first manufacturing method described above.

Next, the second manufacturing method will be described. In this manufacturing method, as shown in FIG. 9, the hydroxyapatite particles 11 are attached to the surface of the amorphous alloy core material 12 through an adhesive or the like. In an argon gas atmosphere having a glass transition temperature Tg or higher and a crystallization temperature Tx or lower, this is set in the molding die 13 and press-pressing 14 is applied from above, whereby a side switch with the inner wall surface of the molding die 13 is performed. After obtaining the state, the hydroxyapatite particles 11 are mechanically embedded in the surface of the amorphous alloy core material 12 to manufacture an artificial dental root.

Under the second manufacturing method, in order to embed the hydroxyapatite particles 11, a press-pressing 1
4 is required, but the pressure atmosphere temperature is 400 to 500 °
C can be used at a relatively low temperature, and the superplasticity of 2000% of the amorphous alloy can also make the pressing force relatively weak. Therefore, compared with the conventional one, the hydroxyapatite particles 11 are crushed carelessly. There is an advantage that it is possible to reduce the occurrence of a gap and to suppress the generation of a gap due to a difference in contraction to a slight extent.

In each manufacturing method, hydroxyapatite, which is the mainstream at present, is used as the biocompatible material, but the present invention is not limited to this, and calcium phosphate-based materials are used. It goes without saying that non-calcium phosphate-based ceramic materials and other biocompatible materials such as bioglass can be used. In particular, the mainstream hydroxyapatite is commercially available in the form of particles, and in order to easily join the commercially available apatite as a sintered body or in the form of particles, it is preferable to use an amorphous alloy for the metal core material. I can say.

[0021]

As described above, according to the present invention, by using an amorphous alloy having a supercooled liquid region for the metal core material, compared with the conventional one using titanium or titanium alloy for the metal core material, in particular For the first time, it has become possible to provide an artificial dental root having excellent bonding strength with a compatible material. Further, in the case of manufacturing an artificial tooth root in which the metal core material is made of an amorphous alloy, it becomes possible to manufacture the artificial tooth root in an extremely simple processing step, which contributes to cost reduction of the artificial tooth root itself.

[Brief description of drawings]

FIG. 1 is an explanatory view showing a method for manufacturing a bulk material as a sample.

FIG. 2 is an explanatory diagram showing a diffraction result by an X-ray diffractometer.

FIG. 3 is an explanatory diagram showing a differential scanning calorimetric analysis result.

FIG. 4 is an explanatory diagram showing a composition ratio of three elements and a composition dependency of a supercooled liquid region.

FIG. 5 is an explanatory diagram showing a stress-strain curve with respect to temperature change.

FIG. 6 is an explanatory view showing a method for manufacturing an amorphous alloy wire.

FIG. 7 is an explanatory view showing a sintered body of a biocompatible material used in the first manufacturing method.

FIG. 8 is an explanatory view showing a state where an amorphous alloy core material is pushed into the sintered body.

FIG. 9 is an explanatory diagram showing a second manufacturing method.

[Explanation of symbols]

 10 Sintered body of hydroxyapatite 11 Particles of hydroxyapatite 12 Amorphous alloy core material 13 Mold 14 Pressing pressure

─────────────────────────────────────────────────── ─── Continuation of the front page (51) Int.Cl. ⁶ Identification code Internal reference number FI technical display location C22C 23/00 45/02 Z 45/04 Z 45/10 (72) Inventor Akihisa Inoue Sendai, Miyagi Prefecture Kawauchi Mubanji, Aoba-ku, Aichi-shi 11-806 (72) Inventor Zhang Bao 1-10-12 Togomi, Aoba-ku, Sendai, Miyagi Large letters Shimono Upper letters Kanayadaira 417-2 (72) Inventor Kikuo Nishi 1-121 Hashimotocho, Haramachi City, Fukushima Prefecture

Claims (7)

[Claims] 1. A dental artificial dental root having a metal core material whose surface is coated with a biocompatible material, wherein the metal core material is made of an amorphous alloy having a supercooled liquid region. 2. The dental artificial tooth root according to claim 1, wherein the supercooled liquid region is 20K or more. 3. The dental artificial dental root according to claim 1, which has an amorphous alloy ratio of 30% or more. 4. A general formula: Xa-Yb-Mc (X is Zr, T
i, Hf, Mg and one or more metals selected from rare earth metals, Y is one or more metals selected from Al, Zr, Hf, Ti and rare earth metals, M is a transition such as Fe, Co, Ni and Cu. One or more metals selected from metals, a = 30
~ 80, b = 5 to 20, c = 0 to 60), the dental artificial dental root according to claim 1 or claim 2 or claim 3 consisting of an amorphous alloy. 5. Zr ₆₃ --Al ₁₂ --Co ₃ --Ni ₇ --C
_{u 15, Zr 60 -Al 15 -Co} 5 -Ni 15 -Cu 5, Zr 65 -Al 7. 5 -Cu 27. 5, Zr 55 -Al 20 -Co 20, Zr 70 -Al 15 -Fe 15, Zr ₆₀ -Al ₁₅ -Ni _25, dental artificial tooth root according to claim 4, wherein an amorphous alloy having a composition selected from among. 6. A method of manufacturing a dental artificial dental root having a metal core material whose surface is coated with a biocompatible material, wherein a sintered body of the biocompatible material is formed into a tubular shape, and the tubular sintering is performed. A manufacturing method for manufacturing an artificial dental root in which a metal core material is made of an amorphous alloy by pushing an amorphous alloy in a supercooled liquid region into a body. 7. A method for producing a dental artificial dental root having a metal core material whose surface is coated with a biocompatible material, wherein the metal core material is formed of an amorphous alloy, and the amorphous alloy core material is in a supercooled liquid region. Manufacture of a dental artificial dental root in which the metal core material is an amorphous alloy by embedding the particles of the biocompatible material in the surface of the amorphous alloy while adhering the particles of the biocompatible material to the surface of the Method.